Search Results for: nanomedicine

Clottocytes: Artificial Mechanical Platelets

By Robert A. Freitas Jr.Research Scientist, Zyvex LLC

Robert A. Freitas Jr.

People often ask for examples of the unique benefits that nanorobots can bring to medicine. That is, what sorts of simple things will robotic nanomedicine allow us to do, that an advanced biotechnology could not accomplish, even in principle? The respirocytes [1] — artificial mechanical red blood cells — are one answer to this perennial question. Respirocytes are micron-sized diamondoid storage tanks for transporting respiratory gases throughout the human body, that can be reversibly pressurized up to 1000 atm in direct response to changing tissue requirements. Here, I’d like to describe another interesting example of a simple nanorobotic application that could provide a unique superbiological capability: “instant” hemostasis using clottocytes, or artificial mechanical platelets.

The structure and primary functions of the platelet are well-known. In brief, platelets are roughly spheroidal nucleus-free blood cells measuring ~2 microns in diameter with an average bloodstream lifetime of ~10 days [2] and a mean blood concentration of ~250,000 cells/mm3 [3]. Platelets gather at a site of bleeding. There they are activated, becoming sticky and clumping together to form a plug that helps seal the blood vessel and stop the bleeding. At the same time, they release substances that help promote clotting. Natural blood coagulation is a complex process involving platelets, red and white cells, endothelial cells, an array of coagulation factors, fibrinolytic proteins and protease inhibitors whose contributions wax and wane over time. Interestingly, it has been found that platelets can slowly crawl across surfaces [4], and they have other well-studied ancillary abilities such as the phagocytosis of foreign particles [5] and the killing of microfilarial larval parasites [6].

A complete functional design of an artificial platelet is beyond the scope of this paper. Here, I want to focus on the purely mechanical aspects of the hemostatic function of platelets, and describe how this function might be served more effectively by a small in vivo population of medical nanorobotic devices.

After injury to a blood vessel, a natural hemostatic plug is formed which is composed predominantly of platelets. Platelet activation — primary hemostasis — normally proceeds in three phases [13]. The first phase is platelet adhesion, in which a cell monolayer carpet forms in response to the exposure of an appropriate surface to the blood. The relevant natural surface in vivo is thought to be the subendothelial matrix. This matrix lies just below the endothelial cells that coat the blood vessels and would become exposed if a vessel was injured. Artificial surfaces can also induce adhesion. The second phase is platelet aggregation into a plug, mediated by the interaction of fibrinogen with glycoprotein receptors on the platelet surface (the gpIIb/IIIa complex) in the presence of micromolar concentrations of calcium. The third phase is platelet secretion, in which platelet granules release their contents into the extracellular space. These contents include adenosine diphosphate (ADP), Ca++, and various proteins such as platelet factor 4 and thromboglobulin that contribute to the formation of a stable plug, along with other agents such as serotonin and epinephrine which cause vasoconstriction.

Secondary hemostasis then ensues with the deposition of fibrin. Fibrin strands quickly form a fine meshwork of random fibrils, trapping more platelets and other blood cells to produce a solid clot. Total bleeding time, as experimentally measured from initial time of injury to cessation of blood flow, may range from 2-5 minutes [7-9] up to 9-10 minutes [10, 13] if even small doses of aspirin are present, with 2-8 minutes being typical in clinical practice; minor prolongations up to 15-20 minutes are not considered clinically risky [11-13], and medical dictionaries [14] give the normal coagulation time as 6-17 minutes (360-1020 sec). (Bleeding times begin to be prolonged in otherwise normal patients when their platelet count falls below ~50,000 cells/mm3 [12], ~20% of the normal concentration.) Over the next several hours, the fibrils slowly diffuse within the clot, much as spaghetti moves in boiling water, forming unstable side-to-side monomer associations and thereafter thick bundles, until finally they become crosslinked with covalent disulfide bonds by factor XIIIa, making a dense clot. Note that bleeding time is a measure only of clotting due to platelet function and does not account well for the effect of fibrin (the actual coagulant cascade).

By contrast, the artificial mechanical platelet or clottocyte may allow complete hemostasis in as little as ~1 second, even in moderately large wounds. This response time is on the order of 100-1000 times faster than the natural system. Our baseline clottocyte is conceived as a serum oxyglucose-powered spherical nanorobot ~2 microns in diameter (~4 micron3 volume) containing a fiber mesh that is compactly folded onboard. Upon command from its control computer, the device promptly unfurls its mesh packet in the immediate vicinity of an injured blood vessel — following, say, a cut through the skin. Soluble thin films coating certain parts of the mesh dissolve upon contact with plasma water, revealing sticky sections (e.g., complementary to blood group antigens unique to red cell surfaces) in desired patterns. Blood cells are immediately trapped in the overlapping artificial nettings released by multiple neighboring activated clottocytes, and bleeding halts at once.

How much netting can each individual clottocyte carry? The required fiber volume of a mesh that covers an area Anet using fibers of working strength sigmafiber and thickness tfiber with a grid size of lmesh is Vmesh=2(Anet1/2+Anet/lmesh) tfiber2. Minimum fiber thickness is tfiber >~ (pbloodlmesh2/4 sigmafiber)(1/2), where the maximum blood pressure that may be resisted by the netting is pblood ~ 0.25 atm (190 mmHg). Taking lmesh ~1 micron and sigmafiber ~1010 N/m2 for diamondoid fibers, tfiber >~0.8 nm and Vmesh ~0.1 micron3 (taking up 3% of device volume) for a net of area Anet = 0.1 mm2. If instead of the strongest diamondoid fibers we use bioresorbable organic fibers for the netting — having, say, roughly the strength of cellulose or spider silk (sigmafiber ~109 N/m2 [3]) — then the fibers must be 2.5 nm thick and mesh volume becomes 1.3 micron3 (occupying 30% of nanodevice volume) to throw out a 0.1 mm2 net. Stokes law drag power [3] on the net during its 1 second unfurlment time is a fairly modest ~100 pW per nanorobot, assuming whole-blood viscosity at the normal ~45% hematocrit (Hct).

How many clottocytes are needed to stop bleeding in ~1 second? The required blood concentration nbot of nanorobots required to stop capillary flow at velocity vcap in a response time tstop, assuming noverlap fully overlapped nets, is nbot ~noverlap/(Anettstopvcap). Taking noverlap=2, Anet=0.1 mm2, tstop=1 sec, and vcap ~1 mm/sec [3] gives nbot=20 mm– 3, or just ~110 million clottocytes in the entire 5.4-liter human body blood volume representing ~11 m2 of total deployable mesh surface. This total dose is ~0.4 mm3 of clottocytes, which produces a negligible serum nanocrit [3] of Nct ~ 0.00001%. During the 1 second hemostasis time, an incision wound measuring 1 cm long and 3 mm deep would lose only ~6 mm3 of blood, less than one-tenth of a single droplet. There are 2-3 red cells per deployed 1 micron2 mesh square, more than enough to ensure that the meshwork will be completely filled, allowing complete blockage of a breach.

Special control protocols are needed to guarantee that clottocytes don’t release their mesh packets in the wrong place inside the body, or at an inappropriate time. These protocols will demand that carefully specified constellations of sensor readings must be observed before device activation is permitted.

For example, the atmospheric concentrations of gases such as carbon dioxide and oxygen are different than in blood serum. As clottocyte-rich blood enters a breach in a blood vessel, nanorobot onboard sensors can rapidly detect the change in partial pressures, indicating that the nanodevice is being bled out of the body. At a nanorobot whole-blood concentration of 20 mm– 3, mean device separation is 370 microns. If the first device to be bled from the body lies 75 microns from the air-serum interface, oxygen molecules from the air can diffuse through serum at human body temperature (310 K) from the interface to the nanodevice surface in ~1 second [3]. Detection of this change can be rapidly broadcast to neighboring clottocytes using acoustic pulses that are received in times on the order of microseconds, allowing rapid propagation of a device-enablement cascade. Similarly, air temperature is normally lower than body temperature. The thermal equilibration time [3] across a distance L in serum at 310 K is tEQ ~(6.7 x 106) L2, hence a device that lies 75 microns from the air-plasma interface can detect a change in temperature in tEQ ~ 40 millisec. Other relevant sensor readings may include blood pressure profiles, bioacoustic monitoring, bioelectrical field measurements, optical and ultraviolet radiation detection, and sudden shifts in pH or other ionic concentrations. At some cost in rapidity of response, clottocytes also could eavesdrop [3] on natural biological platelet control signals, using sensors with receptors for the natural prostaglandins produced by endothelial cells that normally induce or inhibit platelet activation.

The rapid mechanical action of clottocytes could interfere with the much slower natural platelet adhesion and aggregation processes, or disturb the normal equilibrium between the clotting and fibrinolytic systems. Thus it may be necessary for artificial platelets to release quantities of various chemical substances that will encourage the remainder of the coagulation cascade to proceed normally or at an accelerated pace, including timed localized vasodilation and vasoconstriction, control of endothelial cell modulation of natural platelet action, and finally clot retraction and fibrinolysis much later during tertiary hemostasis.

There is a small risk that a potentially-fatal catastrophic clotting cascade called disseminated intravascular coagulation (DIC) [18] could be triggered by excessive clottocyte activity. Coagulation is usually confined to a localized area by a combination of bloodflow, localized thrombin production, and circulating coagulation inhibitors such as antithrombin III (a potent thrombin inhibitor). But if the stimulus to coagulation is too great, excess thrombin is produced and enters the general circulation. This overwhelms the natural control mechanisms and leads to excess fibrin deposition, formation of large numbers of microthrombi (intravascular clotting), rapid depletion of platelets and fibrinogen (and other coagulation factors), secondary fibrinolysis, and often hemorrhage, the typical signs of acute DIC. One solution is to equip clottocytes with sensors to detect decreased serum levels of fibrinogen, plasminogen, alpha2-antiplasmin, antithrombin III, factor VII and protein C, and elevated levels of thrombin and various fibrin/fibrinogen-derived degradation products. If DIC conditions arise, nanorobots might respond by absorbing and metabolizing the excess thrombin, or by releasing thrombin inhibitors such as antithrombin III, hirudin, argatroban or lepirudin [19] or anticoagulants that reduce thrombin generation such as danaparoid [19] to interrupt the cascade. For example, a ~0.02% Nct concentration of nanorobots, suitably activated, could replace the entire depleted natural bloodstream content of antithrombin III from onboard stores.

Extended bleeding often serves to cleanse a wound of foreign matter and bacteria that might have entered the body along with the skin-penetrating object that caused the wound. By immediately staunching the flow, it might be argued that clottocytes would prevent this natural cleaning action from taking place. However, it is anticipated that clottocytes would only be one component of a complete hematological “upgrade” package, and that other species of circulating nanorobots would perform these scavenging and janitorial tasks.

Yet another possible complication is that the bare tissue walls of a wound will continue to exude fluid, and may begin to desiccate, if only the capillary termini are sealed but the rest of the tissue is left exposed to open air. Since clottocytes may remain attached to their discharged nets, and can communicate with each other via acoustic channels [3], it should be possible to precisely control the development of a larger artificial mesh-based clot via coordinated mesh extensions or retractions within the clot. Alternatively, clottocytes could allow blood fluids to flood small incised or avulsed wound volumes, allowing exposed tissue walls to be bathed in fluids but casting a watertight sealant net across the wound opening flush with the epidermal plane of the wound cavity.

What about internal bleeding? Clottocytes will require far more sophisticated operational protocols if they are intended to assist platelets participating in the sealing of internal blood vessel lesions, in order to avoid inadvertently blocking the lumen of the entire vessel. Similarly, prevention of bleeding at vascular anastomoses, hemarthroses, internal bruising, “blood blisters” and larger tissue hematomas, as well as forced local coagulation in tumors or in intracerebral aneurysms, may also require more advanced protocols, possibly including integration with pre-existing in vivo navigation systems [3]. For some of these applications, motile clottocytes may be required in place of the free-floating nanorobots described in this paper, along with a graduated recruitment response depending upon how many (intercommunicating) devices appear to be involved in the event.

Numerous significant design questions remain, including most importantly biocompatibility issues — are clottocytes truly inert? Will they interact with other blood cells or with endothelial cells? Will they activate complement pathways or elicit fibrin deposition? Diamond is indeed chemically inert [15, 16] and is generally considered noninflammatory relative to the complement system [17]. An enveloped clottocyte externally coated with autologous platelet membrane should be nearly as biocompatible as native platelets and could also assist in the recruitment of intrinsic coagulation mechanisms — especially important for severely thrombocytopenic (platelet-poor) patients. Another requirement is that the bioresorbable netting must be capable of being broken up into phagocytosable pieces, either by natural enzymatic pathways or by artificial fiberlytic enzymes (analogous to fibrinolytic plasmin) that may be released from the clottocyte at the appropriate time. The fiber material should also be nonimmunogenic, to avoid uncontrolled immune-mediated platelet activation [19]. Further analysis must await the completion of Volume II of Nanomedicine.

To summarize: an artificial mechanical platelet appears to permit the halting of bleeding 100-1000 times faster than natural hemostasis. While 1-300 platelets might be broken and still be insufficient to initiate a self-perpetuating clotting cascade, even a single clottocyte, upon reliably detecting a blood vessel break, can rapidly communicate this fact to its neighbors, immediately triggering a progressive controlled mesh-release cascade. Clottocytes may perform a clotting function that is equivalent in its essentials to that performed by biological platelets — but at only ~0.01% of the bloodstream concentration of those cells. Hence clottocytes appear to be ~10,000 times more effective as clotting agents than an equal volume of natural platelets.

Acknowledgments

The author thanks Stephen S. Flitman, M.D., C. Christopher Hook, M.D., and Ronald G. Landes, M.D., for helpful comments on an earlier version of this paper.

Will Serum Proteins Stick to Nanorobot Surfaces?

By Robert A. Freitas Jr.Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

When an artificial nanoorgan is implanted in the body, it may be desirable to promote rapid adhesion to cells and tissues. But for medical nanorobots floating or swimming in the circulation such adhesion will not normally be desirable. Thus one preliminary question is whether sticky biological "gunk" will adhere to the surface of a diamondoid nanorobot when it is placed in the bloodstream, and if so, what can be done about it? As the famous physiological chemist Leo Vroman once hyperbolized [1]: "Facing a hail of miscellaneous eggs, we cannot expect to come away clean. Unless they are hard-boiled ones, we are most likely to become coated rapidly with a relatively thin film of matter from the most numerous and most fragile eggs. Similarly, no interfaces may exist that, facing blood plasma, can escape being coated with the most abundant and fragile plasma proteins."

When a foreign material is implanted into a host tissue, the first event to occur at the tissue-material interface which dictates biocompatibility is the noncovalent adsorption of plasma proteins from blood onto the surface [2, 3]. Protein adsorption is much more rapid than the transport of host cells to foreign surfaces. Once proteins have adsorbed to the surface of the foreign material, host cells no longer see the underlying material, but only the protein-coated surface overlayer. This adsorbed protein overlayer — rather than the foreign material itself – then mediates the types of cells that may adhere to the surface, which ultimately can determine the type of tissue that forms in the vicinity [2]. Thus the type and state of adsorbed proteins, including their conformational changes, will be critical determinants of biocompatibility [3-5], including nanorobot-cell interactions and nanorobot surface fouling.

Even by the late 1960s, Vroman and Adams [1], and Baier and Dutton [6] had found that within 10 seconds of exposure to blood or plasma, a uniform ~6 nm layer of fibrinogen formed on surfaces of Ge, Pt, Si, and Ta. After 60 sec, the layer was less uniform and averaged ~12.5 nm thick, but was still dominated by fibrinogen. Rudee and Price [7] determined that human serum albumin (molecular dimensions 8 nm x 3.8 nm) formed a continuous film on amorphous carbon surface in only 1.3 sec of exposure. Fibrinogen required 2.5 sec to form films.

The first direct study of protein adsorption on diamond, done by Tang et al. [3] in 1995, focused on fibrinogen. Fibrinogen, a 340-kilodalton soluble plasma glycoprotein ~47.5 nm in length, is the major surface protein to initiate coagulation [5] (via platelet adhesion to fibrinogen) and inflammation including fibrosis [4] around implanted biomaterials. The adsorption and deformation ("denaturation") of adsorbed fibrinogen molecules is commonly used as a biocompatibility indicator. The amounts of denatured fibrinogen accumulated on surfaces correlates closely with the extent of biomaterial-mediated inflammation [8].

Accordingly, Tang and colleagues [3] measured ~3.7 mg/m2 (~6600 molecules/micron2) gross surface adsorption of human fibrinogen on chemical-vapor-deposited (CVD) diamond surfaces, after incubation of the plasma-coated diamond surface in a 20 microgram/cm3 fibrinogen solution (~0.1% of blood concentration [9]) for 8 hours at room temperature. Much of this adsorbed fibrinogen was only loosely bound, however. A solution of sodium dodecyl sulfate (an anionic detergent or surfactant commonly used to solubilize proteins) was rinsed over the incubated CVD surface to remove the loosely-bound or elutable (non-denatured) fibrinogen and ~48% of the molecules detached, leaving ~2.1 mg/m2 (~3700 molecules/micron2) of spontaneously denatured fibrinogen still present on the CVD diamond surface. These adsorbed amounts are comparable to the amounts adsorbed on titanium and stainless steel materials widely used as surgical implants that are regarded as "biocompatible."

However, CVD diamond might not accurately represent the atomically-smooth flawless diamond surfaces which may characterize the typical MNT-manufactured medical nanorobot exterior. Far from being atomically smooth, CVD diamond films are amorphous and polycrystalline [10], often with grain sizes up to 1-10 microns. In Tang’s experiment, diamond wafers with two distinct sides were tested, as follows:

The nucleation side of the diamond wafers was grown in contact with a flat silicon substrate, which was then dissolved away by acid. The formation of SiC on such a substrate allows silicon to bond well with carbon during the growth process [10]. However, the presence of small amounts of surviving carbide in the nucleation diamond surface, or of concave nanoscale surface features recording the removal of SiC by etchant, could markedly alter the protein adsorbent characteristics of the diamond surface at the molecular level. Also, SiC is tolerated by cells up to 0.1 mg/cm3 concentration but is cytotoxic at 1 mg/cm3 [11]. Furthermore, a contact profilometer measured the nucleation surface as having a rugosity of up to 250 nm, a roughness 100-1000 times greater than that which may be expected at the surface of the typical diamondoid medical nanodevice.

The growth side of the diamond wafers used in Tang’s experiment was much rougher than the nucleation surfaces, so this surface was ground and polished but only to a rugosity of ~1 micron peak-to-valley – roughly the diameter of an entire bloodborne medical nanorobot and clearly not representative of an atomically-precise engineered medical nanodevice surface. There is no indication whether the grinding and polishing of the growth surface was done under oil (thus preserving a predominantly hydrogen-terminated, hence strongly hydrophobic, surface [12]), and there was no evaluation of whether subsequent etching with H2SO4 and H2O2 might have produced carbonyl and hydroxyl conversions at the surface (thus possibly creating regions of hydrophilicity).

Furthermore, diamond crystals are believed to polish by successive repeated microcleavage along {111} planes, which is why polishing is much easier in some directions than others [12]. Non-{111} surfaces, when mechanically polished, will always be rough and will consist of small domains of {111} surface canted at appropriate angles to the macroscopic orientation [12] – residual asperities of ~5 nm have been reported even for extremely carefully polished surfaces. The general conclusion is that the chemical and mechanical processes used in Tang’s experiment seem unlikely to have produced a surface that is well-characterized at the molecular level. Protein adhesion to near-atomically smooth diamond surfaces remains to be investigated experimentally, and may be quite low.

Once the precise molecular mechanisms of protein adhesion are fully understood (this knowledge was still incomplete in 2000), nanodevice surfaces can be altered for maximum proteinophobicity if this is required for a particular medical mission. It is almost certain that such designs are possible because numerous semi-proteinophobic artificial molecular surfaces are already available.

For instance, a polyethylene glycol (PEG) coating on a 200-nm poly(lactic acid) (PLA) nanosphere surface creates a brushlike steric barrier, hindering its opsonization and uptake by the mononuclear phagocyte system, thus increasing its bloodstream half-life [13]. These "pegylated" nanospheres have been investigated as an injectable blood-persistent system for controlled drug release, for site-specific drug delivery, and for medical imaging [14]. Adsorption of human serum albumin (MW = 66 kilodaltons) on pegylated nanosphere surfaces at pH 7.4 at equilibrium (i.e., after 5 days) is 0.15 mg/m2 (~1400 molecules/micron2) compared to 2.2 mg/m2 (~20,000 molecules/micron2) for unpegylated polymer [14], which is of the same order of magnitude as that observed for other unmodified hydrophobic surfaces [15]. Under in vitro conditions at 37o C and pH 7.4, about one-third of the adsorbed PEG detaches from the PLA nanospheres after 2 weeks, at a near-linear detachment rate [16]. Tetraethylene glycol dimethyl ether glow-discharge plasma-deposited surfaces can reduce fibrinogen adsorption to ~0.2 mg/m2 (~350 molecules/micron2) on many different substrates [17].

Another even more effective way to create nonadhesive nanorobot surfaces may be the biomimetic approach. For example, the external region of a cell membrane, known as the glycocalyx, is dominated by glycosylated molecules, which direct specific interactions such as cell-cell recognition and contribute to the steric repulsion that prevents undesirable non-specific adhesion of other molecules and cells. Holland and colleagues [18] have modified a pyrolytic graphite surface by attaching oligosaccharide surfactant polymers which, like a glycocalyx, provides a dense and confluent layer of oligosaccharides that mimics the non-adhesive properties of a glycocalyx. The surfactant polymers consist of a flexible poly(vinyl amine) backbone (MW ~ 6000 daltons, diameter 0.25 nm) with multiple randomly-spaced dextran (MW ~ 1600 daltons, diameter ~0.9 nm) and alkanoyl (hexanoyl or lauroyl) side chains which constrain the polymer backbone to lie parallel to the substrate. Solvated dextran side chains protrude into the aqueous phase, creating a glycocalyx-like monolayer coating 0.7-1.2 nm thick as measured by tapping-mode AFM [18]. Dextran has a stable helical structure; steric repulsion between adjacent dextrans is believed to produce a brush-like conformation. In vitro experiments show that the resulting biomimetic surface, which the authors assert undergoes spontaneous adsorption on diverse hydrophobic surfaces, is effective in suppressing at least ~90% of all plasma protein adsorption from human plasma protein solution [18]. According to the authors:

"The steric barrier provided by the highly hydrated dextrans is designed to suppress non-specific adsorption of plasma proteins [19], whereas the high energy of desorption and low water solubility of the adsorbed surfactant polymer is designed to minimize possible displacement or exchange reactions with highly surface-active plasma proteins."

Diamond particles have already been encapsulated inside stealth liposomes. Stealth liposomes are relatively nonadhesive polyhydroxylated species that exhibit reduced recognition and uptake by the body’s reticuloendothelial system, and give a circulation half-life of ~1 day [20]. In one experiment to test a possible artificial oxygen carrier [21], hemoglobin molecules were irreversibly adsorbed onto carbohydrate-coated diamond particles measuring ~75 nm in diameter, then were encapsulated in a standard mixture of phospholipids, yielding preparations of spherical liposomes which were stable for >48 hours with bound-Hb concentrations near 100 gm/liter with as little as 1% free Hb. Evaluation of oxygen lability showed normal sigmoidal O2 binding behavior with p50 from 12 mmHg up to 37 mmHg under control of an allosteric effector [21].

Whether pure atomically-smooth diamondoid materials will give us sufficiently nonadhesive surfaces, or if instead thin engineered coatings or active semaphoric surfaces [9] will be necessary to ensure adequate biocompatibility of medical nanorobots, is an outstanding research issue that can best be resolved by future experiments. This is a very critical topic because, unlike the materials used in a joint prosthesis, nanorobots will be present in the microvasculature of critical organs. The adhesiveness of many hundreds of serum proteins to the artificial nanorobot calyx must be assayed, and the relative serum concentration of these proteins changes according to the time of day or the physiological state of the individual (e.g., TNF, IL-1, IL-2, and transferrin rise dramatically in the acute phase response to a pathogen). Relevant investigations are to be encouraged at the earliest possible opportunity. The author welcomes and invites further dialogue on this subject by interested specialists.

Nanopyrexia

By Robert A. Freitas Jr.Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

When considering the biocompatibility of medical nanorobots [1], an important concern is whether such devices may inadvertently act as “pyrogens”  agents causing systemic body temperature suddenly to rise, mimicking the effects of an infection. To understand this potential medical risk and its possible resolution, we must first examine how body temperature is normally controlled.

Human core temperature is tightly regulated through the preoptic nucleus of the anterior hypothalamus [2] to a mean “set point” of 37o C with circadian variations around this mean rarely exceeding 0.6o C. An array of thermoregulatory mechanisms ensures that the hypothalamic thermal set point temperature is maintained to within a natural “load error” of 0.2-0.5o C. Thermal deviations exceeding the load error provoke a natural counteractive response to restore core temperature back to the set point.

Abnormal elevation of systemic body temperature (called “pyrexia”) can occur in one of two ways: hyperthermia or fever [3].

In hyperthermia, thermal control mechanisms are overwhelmed, so that heat production exceeds heat dissipation. Hyperthermia may develop during periods of intense physical exertion, dehydration, immersion in hot fluids, or from waste heat thrown off by energy-consuming nanorobots in vivo [1]. In each case the body’s thermoregulatory mechanisms are fully engaged, attempting to cope with the departure from homeostasis. In some situations, thermoregulatory disorders such as heatstroke, hypothalamic insult (caused by drugs, infection or tumor), malignant hyperthermia of anesthesia, or thyroid storm, can cause extreme pyrexia with temperature rising to 41.1o C or higher. Heating blood above 47o C rapidly produces visible damage to erythrocytes [4]; heat-damaged cells show morphologic changes, increases in osmotic and mechanical fragility, and are removed rapidly after reinjection into the circulation. Similarly, an increase of ~6.5-10o C in tooth pulp temperature for > 30-45 seconds (e.g., due to overactive dental nanorobots [5]) can permanently damage the pulp [6]. If nanorobots are the cause of hyperthermia, it is because local or systemic thermogenic limits [1] are being exceeded. Obeying these operational limits should avoid the problem.

In fever, the second cause of pyrexia and the primary topic of this paper, the hypothalamic thermal set point is shifted higher. Fever is a natural self-defense mechanism (produced by substrate cycling in skeletal muscle) intended to make the host less hospitable to microscopic invaders. The intact control mechanisms of thermoregulation act to raise body temperature up to the new set point, then maintain the elevated systemic temperature. Thus fever is not equivalent to an elevated core temperature, but rather represents an elevated set point [7]. Fever is triggered by the release of endogenous pyrogen (a fever-producing substance) from cells of the immune system into the bloodstream. Mononuclear phagocytes are the main source of endogenous pyrogens, and a variety of these substances, often categorized as monokines and lymphokines, or collectively, as cytokines, also mediate the acute-phase response to infection and inflammation. Pyrogenic cytokines act as hormones in that they are carried by the circulation from the local inflammatory site of production to the central nervous system, where they bind with high affinity to 80 kDa receptors present on vascular endothelial cells within the hypothalamus. This elicits phospholipases which in turn cause release of arachidonic acids from membrane phospholipids, as a result of which prostaglandin levels rise, resetting the hypothalamic regulatory center to a new set point; the cytokines may also interact directly with neural tissues [7].

IL-1 (17.4 kDa) comes mainly from monocytes and macrophages, though it can also be produced by neutrophils, B and T cells, endothelial cells, and virtually all other nucleated cells [3]. IL-1 production may be stimulated by the presence of microorganisms, exposure to endotoxin and other bacterial toxins or microbial products, phagocytosis, antigen-antibody immune complexes, and various forms of tissue injury [3], and IL-1 induces additional IL-1 production and additional IL-1 receptor expression on certain target cells. IL-1 also stimulates immune cells thus enhancing host defense mechanisms, stimulates lactoferrin release by neutrophils (which have ~1700 IL-1 receptors per cell) reducing serum iron levels during many bacterial infections thus retarding bacterial growth, acts on the central nervous system to induce sleep, and has numerous other biologic properties.

TNF is another pyrogenic cytokine that acts directly on the hypothalamus to elevate the thermal set point, and also causes fever by stimulating IL-1 production. Macrophages are the main source of TNF, along with monocytes and NK cells. TNF production is stimulated most potently by endotoxin, but also by certain parasites, viruses, enterotoxins (including toxic-shock syndrome toxin-1), and IL-1. Peak serum levels occur in 90 minutes, but TNF is cleared from the circulation in ~3 hours [3]. TNF binds to different receptors than IL-1  these are found in the CNS, on vascular endothelium, adipose tissue, and on liver, kidney and lung tissues [3]. TNF has other biological properties besides pyrogenicity, including increasing resistance to infection, inhibition of ACTH release, induction of sleep, and mediation of septic shock.

Can nanorobots act as pyrogens, inducing systemic fever (nanopyrexia)? Certainly, any external organic coatings on nanorobots should be verified as nonpyrogenic. For example, phagocytosed latex particles do not stimulate pyrogen production in macrophages [8]. Fever occurs in about one-third of all hospital patients, 67% of these due to infection but 12%-18% due to “fever of unknown origin” or FUO that is nonetheless almost certainly biochemically mediated. FUO is usually ascribed to infections, neoplasms, collagen vascular disease, granulomatous diseases (including starch peritonitis, a febrile granulomatous response to starch introduced on surgical gloves), chronic liver disease and IBD, pulmonary emboli and atelectasis, and very rarely certain drugs such as Dilantin. Fever can also be produced by mechanical tissue disturbance such as a thoracic esophageal perforation [9], excision of Teflon particulate masses [10], knee and hip arthroplasty, or shock wave lithotripsy [11], confirming the need for cautious nanosurgery.

As of 2000, there are no reports of pyrogenicity for anticipated nanorobot simple building materials such as diamond, fullerenes, or graphite. Carbon powder has been used in nasal provocation tests without eliciting fever, though there are rare cases of fever from amorphous carbon particles in India ink [12]. With rare exception, bulk Teflon appears nonpyrogenic in vivo, although perfluorocarbon emulsion can cause cutaneous flushing and fever at low doses and “polymer fume fever” [13] or “Teflon fever” results when Teflon combustion products are inhaled.

No pyrogenicity of monocrystal sapphire has been reported. However, there is one case of fever possibly caused by alumina powder inhalation [14]. Additionally, while ceramics appear generally nonpyrogenic, macrophages exposed to particulate alumina ceramic release TNF, increasingly with size and concentration of particles [15].

Other particulates are less inert. Metal fume fever (due to zinc oxide inhalation) is well-known and excess trace elements such as copper and zinc can induce fever. Phagocytosed silica crystals do elicit pyrogen [16, 17] and various low-solubility substances that crystallize in the human body can trigger fever once the crystals have formed. For example, monosodium urate monohydrate crystals [16-18] which are deposited in synovial fluid during gout, causing fever, stimulate IL-1, TNF, and IL-6 production in monocytes or synoviocytes, with smaller 10-40 micron crystals less pyrogenic than the larger aggregates [16]. Calcium pyrophosphate dihydrate (CPPD) deposited in the fibrocartilage during chondrocalcinosis (aka CPPD crystal deposition disease) is pyrogenic [19], and CPPD crystals increase IL-6 production by monocytes and synoviocytes in vitro [18]. Fever has been reported from nephrolithiasis (kidney stones), from crystalluria with calcium oxalate or phosphate in urolithiasis (bladder stones), from calcified lymph-node stones in broncholithiasis, from calcified salivary gland stones in sialolithiasis, and from precipitated crystals in the pulmonary microvasculature in a patient receiving total parenteral nutrition. Cholesterol crystals deposited as gallstones during cholelithiasis may be pyrogenic, as are cholesterol crystal emboli in the blood [20]. A systematic assessment of pyrogenicity should be undertaken for all crystalline and ceramic materials likely to be employed in the construction of medical nanorobots.

If inherent nanodevice surface pyrogenicity cannot be avoided, the pyrogenic pathway is readily controlled by in vivo medical nanorobots because only a small number of critical mediators are involved. For instance, the cytokine IL-4 suppresses production of the endogenous pyrogens IL-1, TNF and IL-6 [21], and NSAID prostaglandin inhibitors like aspirin or ibuprofen are also effective antipyretic agents that block prostaglandin synthase (cyclooxygenase) enzyme activity and thus block prostaglandin production. Antagonists of the IL-1 receptor have been identified [22]. Glucocorticoids inhibit the production of IL-1, TNF and IL-6, and other inhibitors of TNF are known such as the anti-TNF monoclonal antibody Etanercept currently used in rheumatoid arthritis patients with excellent results. Nonsteroidal anti-inflammatory antipyretic drugs are used for treatment of gout and other crystal-induced arthropathies [23]. Nanorobots may release these (or similar) inhibitors, antagonists or down-regulators in a targeted fashion to interrupt the pyrogenic pathway, or may use molecular rotors to selectively absorb the endogenous pyrogens, chemically modify them, then release them back into the body in a harmless inactivated form.

For example, typical bloodstream concentrations are ~10 pg/cm3 for IL-1 beta [24] and ~100 pg/cm3 for TNF [15], or ~0.0003-0.003 molecules/micron3 assuming a molecular weight of ~17.4 kDa for either molecule [7]. If there are 2-20 x 1012 molecules of these cytokines in the entire circulation, then a fleet of 0.1-1 trillion nanorobots each with 10,000 sorting rotors on its surface (extracting ~0.0001 molecules/rotor-sec [1]) can reduce bloodstream IL-1 or TNF concentrations by ~99% in ~20 seconds. Selective absorption of prostaglandins, present in blood plasma at ~400 pg/cm3 [1], might also serve to “manually” reduce the hypothalamic thermal set point. One other possible approach, adopted by certain vaccinia virus strains [25], is to suppress the fever response by releasing soluble IL-1 receptors that bind to IL-1, thus inhibiting this normal pathway.

It is possible that perfectly biocompatible-surfaced nanorobots cannot be designed, or that necessary additional anti-pyrogenic functions cannot be added to nanorobotic devices already hard-pressed for onboard space. Although not ideal, in such cases a collection of different nanodevices could be deployed to implement a given treatment. Some devices would attend to the primary therapeutic goal with others attending to the management of the unwanted biological responses, crudely analogous to drug combinations in current medical practice such as demerol plus vistaril or combinations of chemotherapeutics and anti-emetics.

Acknowledgments

The author thanks Stephen S. Flitman, M.D., and C. Christopher Hook, M.D., for helpful comments on an earlier version of this paper.

Microbivores: Artificial Mechanical Phagocytes

By Robert A. Freitas Jr.Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

Nanomedicine [1] offers the prospect of powerful new tools for the treatment of human diseases and the augmentation of human biological systems. Diamondoid-based medical nanorobotics may offer substantial improvements in capabilities over natural biological systems, exceeding even the improvements possible via tissue engineering and biotechnology. For example, the respirocytes [2]  artificial red blood cells comprised of microscopic diamondoid pressure tanks that are operated at up to 1000 atm of pressure  could carry >200 times more respiratory gases than an equal volume of natural red blood cells. The clottocytes [3] are artificial platelets that could stop human bleeding within ~1 second of physical injury, but using only 0.01% the bloodstream concentration of natural platelets  in other words, nanorobotic clottocytes would be ~10,000 times more effective as clotting agents than an equal volume of natural platelets.

In this article I’d like to examine the future nanorobotic equivalent of the third major class of natural blood cells  the white cells. This paper summarizes the results of a recently completed scaling study of artificial mechanical phagocytes of microscopic size, called “microbivores.” Microbivores constitute a potentially large class of medical nanorobots intended to be deployed in human patients for a wide variety of antimicrobial therapeutic purposes, for example as a first-line response to septicemia. The analysis focuses on a relatively simple device: an intravenous (I.V.) microbivore whose primary function is to destroy microbiological pathogens found in the human bloodstream, using the “digest and discharge” protocol first described by the author elsewhere [1]. The full technical paper describing the microbivore scaling design study is already available online [4].

Septicemia, also known as blood poisoning, is the presence of pathogenic microorganisms in the blood. If allowed to progress, these microorganisms can multiply and cause an overwhelming infection. Bacteremia is the presence of bacteria in the human bloodstream. Although bacterial nutrients are plentiful in blood, the healthy human bloodstream is generally considered a sterile environment. Major antimicrobial defenses include the circulating neutrophils and monocytes (white cells) capable of phagocytosis (engulfing and digesting other cells) and the supporting components of humoral immunity including complement and immunoglobulins.

Still, it is not unusual to find a few bacteria in blood. Normal activities like chewing, brushing or flossing teeth causes movement of teeth in their sockets, infusing a burst of commensal oral microbes into the bloodstream [5]. Bacteria can enter the blood via an injury to the skin, the lining of the mouth or gums, or from gingivitis and other minor infections in the skin and elsewhere [6]. Bacteria can also enter the blood during surgical, dental, or other medical procedures such as the insertion of I.V. lines (providing fluids, nutrition or medications), cystoscopy (a viewing tube inserted to examine the bladder), colonoscopy (a viewing tube inserted to view the colon), or heart valve replacement with a prosthetic [6]. Such bacteria are normally removed by circulating leukocytes (along with fixed reticuloendothelial phagocytes in the spleen, liver, and lungs), but a few species of bacteria are unusually virulent and can overwhelm the natural defenses. The Center for Disease Control estimates that ~25,000 U.S. patients die each year from bacterial sepsis. Current therapies often involve multiple antibiotics administered simultaneously in multi-gram quantities per day. These treatments can sometimes take weeks or even months to bring under control certain hardy infectious microorganisms like Pseudomonas aeruginosa or enterobacteria such as Escherichia coli and Salmonella.

A nanorobotic device that could safely provide quick and complete eradication of bloodborne pathogens using relatively low doses of devices would be a welcome addition to the physician’s therapeutic armamentarium. Such a machine is the microbivore, an artificial mechanical phagocyte.

The microbivore is an oblate spheroidal nanomedical device consisting of 610 billion precisely arranged structural atoms plus another ~150 billion mostly gas or water molecules when fully loaded. The nanorobot measures 3.4 microns in diameter along its major axis and 2.0 microns in diameter along its minor axis, thus ensuring ready passage through even the narrowest of human capillaries which are ~4 microns in diameter [1]. Its gross geometric volume of 12.1056 micron3 includes two normally empty internal materials processing chambers totalling 4 micron3 in displaced volume. The nanodevice consumes 100-200 pW of continuous power while in operation and can completely digest trapped microbes at a maximum throughput of 2 micron3 per 30-second cycle, large enough to internalize a single microbe from virtually any major bacteremic species in a single gulp. As in previous designs [2], to help ensure high reliability the microbivore has tenfold redundancy in all major components, excluding only the largest passive structural elements. The microbivore has a dry mass of 12.2 picograms.

Here’s how the nanorobot works. During each cycle of operation, the target bacterium is bound to the surface of the microbivore like a fly on flypaper, via species-specific reversible binding sites [1]. Telescoping robotic grapples emerge from silos in the device surface, establish secure anchorage to the microbe’s plasma membrane, then transport the pathogen to the ingestion port at the front of the device where the pathogen cell is internalized into a 2 micron3 morcellation chamber. After sufficient mechanical mincing, the morcellated remains of the cell are pistoned into a 2 micron3 digestion chamber where a preprogrammed sequence of 40 engineered enzymes are successively injected and extracted six times, progressively reducing the morcellate ultimately to monoresidue amino acids, mononucleotides, glycerol, free fatty acids and simple sugars. These simple molecules are then harmlessly discharged back into the bloodstream through an exhaust port at the rear of the device, completing the 30-second digestion cycle. This “digest and discharge” protocol [1] is conceptually similar to the internalization and digestion process practiced by natural phagocytes, except that the artificial process should be much faster and cleaner. For example, it is well-known that macrophages release biologically active compounds during bacteriophagy [7], whereas well-designed microbivores need only release biologically inactive effluent.

Natural phagocytic cells are 100-1000 times larger in volume than microbivores but may consume almost as much power during comparable activities. For instance, heat production rises from 9 pW in unstimulated human neutrophils up to 28 pW during phagocytosis, with the rise proportional to the number of particles ingested [8]. The basal rate for a resting ~400 micron3 T-cell lymphocyte is ~20 pW, rising to ~65 pW during antigen response [9].

Microbe ingestion times for natural professional phagocytes can be quite rapid, often a matter of minutes, but full digestion and excretion of the target pathogen may require hours. While macrophages can ingest up to ~25% of their volume per hour [10], microbivores can process ~2000% of their volume per hour, thus are ~80 times more efficient as phagocytic agents. In other words, a given volume of microbivores can digest bacterial pathogens 80 times faster than an equal volume of white cells or macrophages could digest them.

Many natural professional phagocytic cells such as neutrophils also have a maximum capacity for phagocytosis during their short lifetime, typically a few hours in blood or a few days in tissue. In one experiment [11], 1-100 S. aureus or S. faecalis bacteria were presented to each neutrophil, which digested more of them at the higher concentrations. At the highest concentration (100:1), neutrophils from normal patients could only kill a mean of 9 S. aureus bacteria per phagocyte, while neutrophils from carriers of of chronic granulomatous disease could kill a mean of 14 S. faecalis bacteria per phagocyte. By comparison, a single microbivore could completely digest up to ~3000 microbes/day of P. aeruginosa bacteria with no well-defined maximum lifetime capacity for phagocytosis.

In the accompanying technical paper [4] a simple mathematical model for microbivore pharmacokinetics quantifies the activity of a specific dose of nanorobots injected into the human bloodstream, with the conclusion that a 1-terabot (1012-device) dose of microbivores employed in the treatment of a mild bacteremia (0.1 x 106 colony-forming units (CFU) per ml) can reduce the initial whole-bloodstream bacterial load of 5.4 x 108 CFU down to <1 CFU in 460-5400 sec (8-90 min), if 1-10 bacterium-microbivore collisions are required for the bacterium to stick.

Similarly, a severe bacteremia (100 x 106 CFU/ml) is eliminated in 620-7300 sec (10-120 min). Note that a single 1-terabot intravenous dose (a ~12 cm3 injection) of microbivores produces a nanocrit of Nct ~ 0.2% in the blood of a normal adult human male patient and could liberate up to 100-200 watts of systemic waste heat which is very near the maximum thermogenic limit for in vivo medical nanorobot systems [1].

While microbivores can fully eliminate septicemic infections in minutes to hours, natural phagocytic defenses  even when aided by antibiotics  can often require weeks or months to achieve complete clearance of target bacteria from the bloodstream. Thus microbivores appear to be up to ~1000 times faster-acting than either unaided natural or antibiotic-assisted biological phagocytic defenses.

Another useful comparative perspective is that the administration of antibacterial agents (e.g., against E. coli) typically may increase the LD50 of that pathogen by ~500-fold using antibiotics [12] or ~850-fold using monoclonal antibodies [13]. For example, the mammalian LD50 for E. coli is ~0.1-1 x 106 CFU/ml, rising to ~108 CFU/ml with the administration of antibiotics. By employing a suitable dose of microbivores, a bloodstream bacterial concentration up to the theoretical maximum of ~1011 CFU/ml (~20% of blood volume assuming ~2 micron3 organisms) could be controlled, bringing another ~1000-fold improvement using nanomedicine and at last extending the therapeutic competence of the physician to the entire range of potential bacterial threats, including locally dense infections.

With minor additions to the basic design, microbivores could be used to combat toxemia, the distribution throughout the body of poisonous products of bacteria growing in a focal or local site, and other biochemical sequelae of sepsis. For instance, E. coli-induced septicemic shock in vervet monkeys occurred at 425 x 106 CFU/ml and bacterial lipopolysaccharide (LPS) endotoxin rose from normal at 0.076 ng/ml to a maximum of 1.130 ng/ml blood concentration [14]. In another study, endotoxin levels during a gram-negative bacterial infection rose from 0.2 to 2 ng/ml in pig blood [15]. Eliminating a bloodstream concentration of ~2 ng/ml of ~8 kDa LPS endotoxin [16] would require the extraction and enzymatic digestion of ~8 x 1014 LPS molecules from the ~5400 cm3 human blood compartment, a mere ~800 LPS molecules per nanorobot assuming a 1-terabot dose of modified microbivores.

The high mortality (up to 30%-50%) associated with gram-negative sepsis is due in large measure to the patient’s reaction to LPS, an endotoxin which induces the production of cytokines such as IL-1beta and IL-6 which leads to an uncontrolled inflammatory reaction resulting in tissue damage and organ failure [17]. We’ve already noted that small quantities (~ng/ml) of LPS are released by living and growing bacteria, but the killing of bacteria using traditional antibiotic regimens often liberates large quantities of additional LPS, potentially up to ~105 ng/ml [17]. Such massive releases as occur with the use of antibiotics will not accompany the use of microbivores, because all bacterial components (including all cell-wall LPS) are internalized and fully digested into harmless nonantigenic molecules prior to discharge from the device. Microbivores thus represent a complete antimicrobial therapy without increasing the risk of sepsis or septic shock.

If the patient presents with a septic condition before the microbivores are introduced, a substantial preexisting concentration of inflammatory cytokines will likely be present and must be extracted from the blood in concert with the primary antibacterial microbivore treatment. All unwanted cytokine molecules may be rapidly and systemically extracted from the blood using a modest dose of respirocyte-class nanodevices [2] such as pharmacytes [1], a combination-treatment approach previously suggested elsewhere [1, 18]. Specifically, a 1-terabot intravenous dose of micron-size pharmacytes [1] each having ~105 cytokine-specific molecular sorting rotors and ~0.5 micron3 of onboard storage capacity could reduce the blood concentration of ~20 kDa IL-1beta and IL-6 cytokines from LPS-elevated levels of ~100 ng/ml [210] (~3 x 10– 9 molecules/nm3) down to normal serum levels of ~10 pg/ml [211] (~3 x 10– 13 molecules/nm3) after only ~200 sec of diffusion-limited pumping, using just ~0.1% of the available onboard storage volume. (Extracting an additional ~105 ng/ml of LPS from the bloodstream would take a similar amount of time and use ~100% of the available onboard storage volume.)

Microbivores could also be useful for treating infections of the meninges or the cerebrospinal fluid (CSF) and respiratory diseases involving the presence of bacteria in the lungs or sputum, and could also digest bacterial biofilms. These handy nanorobots could quickly rid the blood of nonbacterial pathogens such as viruses (viremia), fungus cells (fungemia), or parasites (parasitemia). Outside the body, microbivore derivatives could help clean up biohazards, toxic biochemicals or other environmental organic materials spills, as in bioremediation.

Acknowledgements

The author thanks C. Christopher Hook, M.D., Stephen S. Flitman, M.D., Ronald G. Landes, M.D., and also Forrest Bishop, Robert J. Bradbury, and Ralph C. Merkle, for helpful comments on the technical paper from which this summary article has been abstracted.

How Nanorobots Can Avoid Phagocytosis by White Cells, Part I

By Robert A. Freitas Jr.Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

Any invading microbe that readily attracts white cells (phagocytes) capable of eating it, and then allows itself easily to be ingested and killed, is generally unsuccessful as a parasite. That’s why most successful bacteria interfere to some extent with the activities of phagocytes or find some way to avoid their attention [1]. Bacterial pathogens have devised numerous diverse strategies to avoid phagocytic engulfment and killing, mostly aimed at blocking one or more of the steps in phagocytosis, thereby halting the process [1].

Similarly, natural phagocytic cells presented with any significant concentration of medical nanorobots [2] also may attempt to internalize these nanorobots. How often will such an opportunity arise? There may be an average of one ~730 micron3 granulocyte (e.g., neutrophil) in every ~3 x 105 micron3 of human blood, one ~1525 micron3 monocyte in every ~2 x 106 micron3 of blood, and one >1525 micron3 macrophage in every ~2 x 105 micron3 of human tissues. By random thermal motions in a quiet fluid, a 2-micron nanorobot would trace out a volume containing one neutrophil in ~70 sec at 310 K ([2], Eqn. 3.1), or would diffuse the ~40 micron mean free distance ([2], Eqn. 9.72) between nanorobot and the nearest macrophage in quiet watery tissue in ~4000 sec ([2], Eqn. 3.1). In a small (1 mm diameter) artery with blood flowing at 100 mm/sec, each 2-micron nanorobot, in a total bloodstream population of 1012 such nanorobots, would collide with a white cell once every ~3 seconds near the periphery of the vessel but only once every ~300 seconds near the center of the vessel ([2], Section 9.4.2.2), a rheological disparity that will be amplified by phagocyte margination ([2], Section 9.4.1.3). Studies of macrophage particle-ingestion kinetics show that the number of particles ingested by each phagocytic cell may rise tenfold as the local particle concentration rises from 5 particles per cell to 150 particles per cell [3].

From these crude estimates, it becomes apparent that virtually every medical nanorobot placed inside the human body will encounter phagocytic cells many times during its mission. Thus all nanorobots which are of a size capable of ingestion by phagocytic cells must incorporate physical mechanisms and operational protocols for avoiding and escaping from phagocytes. Ingestion may require from many tens of seconds to half an hour to go to completion, depending upon the size of the internalized particle, so medical nanorobots should have plenty of time to detect and to actively prevent this process. The initial strategy for medical nanorobots is first to avoid phagocytic contact or recognition, and if this fails, then to avoid nanorobot binding to the phagocyte surface, and phagocytic activation.

One simple avoidance method employed by a few pathogens that may occasionally be practical for medical nanorobots is to confine activities to regions of the human body that are inaccessible to phagocytes. For example, certain internal tissues such as the lumens of glands, the urinary bladder and kidney tubules, and various surface tissues such as the skin are not regularly patrolled by phagocytes [1]. The heart and muscle tissues also are relatively macrophage-poor. If reliable methods can be found for the remote (noncontact) detection of nearby phagocytes, akin to the detectability of bacterial metabolic chemical plumes ([2], Section 8.4.3), then most motile nanorobots should be able to outrun any “pursuing” phagocytes.

If remote phagocyte detection methods cannot be made reliably available, and for nonmotile nanorobots, other contact avoidance techniques must be employed. One potentially useful approach is to make use of the natural mediators of cellular chemotaxis (movement along a spatial gradient or directed cell locomotion) and chemokinesis (general random movement or nondirected cell locomotion) [4]. Specific chemicals are known to be chemorepellents, chemotaxis antagonists, chemotactic factor enzymes or antibodies, or negative chemokinesis agents for various cell types.

For example, monocyte migratory inhibition factor inhibits macrophage migration, with a maximum inhibitory effect at 1 ng/ml for both unchallenged and particle-challenged macrophages [5]. Excess zinc immobilizes macrophages [6], and mononuclear cells cultured from hyperimmunoglobulin-E (HIE) patients produced a ~61 kD protein factor that nontoxically inhibited normal neutrophil and monocyte chemotaxis [7] while serum from those patients contained a 30-40 kD inhibitor of granulocyte and monocyte chemotaxis [8]. Phospholipase A2 inhibitors and a ubiquitin-like peptide [9] inhibit neutrophil chemotaxis, leukocyte-specific protein 1 (LSP1) is a negative regulator of neutrophil chemotaxis [10], and polyamines such as putrescine at 1 mM and spermidine at 0.1-0.5 mM inhibit chemotaxis (but not phagocytosis or engulfment) by neutrophils in vitro [11]. Granulocyte locomotion is also inhibited by diclofenac sodium, a nonsteroidal anti-inflammatory agent, at concentrations below 10 micrograms/ml [4], and eicosapentaenoic acid somewhat rigidifies the plasma membrane of human neutrophils, leading to reduced chemotaxis [12]. In other experiments, chemotaxis by human neutrophils toward several common chemoattractants was inhibited by 80%-95%, maximally at a concentration of ~50 microM of the protein kinase inhibitor 1-(5-isoquinolinesulfonyl) piperazine, without affecting the random migration of these white cells [13].

Much phagocyte chemorepellent research occurs in the context of elucidating bacterial avoidance strategies  strategies that might be mimicked by medical nanorobots. Some bacteria or their products inhibit phagocyte chemotaxis. For example, Streptococcal streptolysin O (which also kills phagocytes) is a chemotactic repellent [1], even in very low concentrations. Staphylococcus aureus produces toxins that inhibit the movement of phagocytes; granulocytes are almost immobilized when administered 12 micrograms/ml of purified S. aureus lipase [14]. Pertussis toxin, produced by the bacterium Bordetella pertussis, inhibits chemotaxis of neutrophils and other phagocytes; a PMN-inhibitory factor (PIF) extracted from B. pertussis cells showed little cytotoxicity and inhibited chemotaxis of neutrophils [15]. Fractions of Mycobacterium tuberculosis inhibit leukocyte migration [1], the Clostridium perfringens phi toxin inhibits neutrophil chemotaxis [1], and other “specific antigen” can suppress basophil chemotaxis. Phagocyte chemotaxis is generally reduced by antibiotics such as cefotazime, rifampin, and teicoplanin [16]. Rifampin and tetracyclines inhibit granulocyte chemotactic activity. Leukocyte, lymphocyte and monocyte chemotaxis is inhibited by methylprednisolone and azathioprine, whereas only lymphocytes are chemotactically inhibited by cyclosporine. More research is required to select, or more likely to design, the ideal chemorepellent agent that might be secreted (perhaps at nM concentrations, ~1 molecule/micron3, or less) by, or surface-tethered to, medical nanorobots seeking to avoid contact with phagocytes. Note that bioactive substances released locally by nanorobots can later be retrieved by similar means, thus avoiding nonlocal accumulations of these substances during nanomedical treatment.

Chemorepulsion is adequate for a few devices on simple missions of limited duration, but large numbers of medical nanorobots on longer more complex missions will inevitably come into physical contact with a phagocyte. The least disruption to normal immune processes is achieved if the nanorobot surface can deny recognition to the inquiring phagocyte at the moment of physical contact. Surface-bound moieties are generally preferable to free-released molecules when large numbers of in vivo nanorobots are involved. For example, each nanorobotic member of an internal communication network ([2], Section 7.3.2), stationed perhaps ~100 microns apart throughout the tissues, must continuously avoid being ingested by passing phagocytes. An approach that relies primarily on antiphagocytic chemical releases risks extinguishing all phagocytic activity throughout the body, severely compromising the natural immune system.

By 2001, “long-circulating” phagocytosis-resistant particles [17] and stealth drug carriers [18] have become the objects of active and extensive research. It is well-known that nanoparticle adsorption and internalization by phagocytes are inhibited by the presence of a coating of polysaccharide (e.g., heparin or dextran) chains in a brush-like configuration, or by very hydrophilic coatings. Low phagocytic uptake is achieved using a surface concentration of 2%-5% by weight of PEG, giving efficient steric stabilization (e.g., a distance of ~1.5 nm between two adjacent terminally-attached PEG chains in the covering brush) and avoiding uptake by neutrophils [19]. Experiments by Davis and Illum [18] suggest that polystyrene particles sterically stabilized with adsorbed poloxamer polymer could achieve an extrapolated zero phagocytic uptake using a ~10 nm thick coating on 60 nm diameter particles or a ~23 nm thick coating for 5.25 micron diameter particles, thus eliminating nonspecific phagocytosis. Another study found that pegylated sheep red blood cells (RBCs) were ineffectively phagocytosed by human monocytes, unlike untreated sheep RBCs. Electrical characteristics also are important. Phagocytosis of polystyrene beads (as measured by cellular oxygen consumption) appears strongly dependent on surface potential and thus upon fixed surface charge, and surface charge heterogeneity across domains as small as 1-4 microns can greatly affect phagocytic ability.

Rather than coatings which phagocytes cannot recognize at all, medical nanorobots alternatively could carry surfaces that phagocytes will recognize as “friendly.” For example, coatings that mimic natural immune-privileged cells could be used. Nanorobot exteriors could be covalently bound with essential erythrocyte coat components — a simulated RBC surface could be useful in the bloodstream, but might provoke a response in the tissues. Similarly, fibroblast-like surface might be useful in the tissues, but is not normally seen in the bloodstream and phagocytes might respond to its presence there. Simulated neutrophil or monocyte surfaces would be better, since these cells normally migrate from blood to tissues, hence the immune system expects to see these surfaces virtually everywhere; lymphocytes are likewise normally present in both blood and tissues but are also adept at passing through the endothelial lining, the lymphatic processes, and the lymph nodes without being detained or trapped, eventually returning to the arterial circulation. The ideal solution may be for the medical nanorobot to display a specific set of self-markers at its surface, perhaps including moieties such as CD47. CD47 is a surface protein present on almost every cell type that provides an explicit phagocytic inhibitor signal to NK cells and to macrophages [20].

Microbial pathogens employ similar strategies to create antiphagocytic surfaces that avoid provoking an overwhelming inflammatory response, thus preventing the host from focusing the phagocytic defenses [1]. Enveloped viruses and some bacterial pathogens can cover their external cell surface with components that are seen as “self” by the host’s phagocytes and immune system, a strategy that hides the true antigenic surface. Phagocytes then cannot recognize the bacteria upon contact and the possibility of opsonization by antibodies to enhance phagocytosis is minimized [1]. For example, Group A streptococci can synthesize a capsule composed of hyaluronic acid, the “ground substance” (tissue cement) found in human connective tissue. The streptococcal hyaluronic acid capsule is nonantigenic and thus very effective in preventing attachment of the organism to the macrophage [21]. Additionally, the cytoplasmic membrane of Streptococcus pyogenes contains antigens similar to those found on human cardiac, skeletal and smooth muscle cells, on heart valve fibroblasts, and in neuronal tissues, resulting in molecular mimicry and an immune tolerance response by the host [22]. Other examples include pathogenic Staphylococcus aureus that produces cell-bound coagulase which clots fibrin on the bacterial surface [1], the syphilitic agent Treponema pallidum that binds human fibronectin to its surface [1], and a variety of bacteria that cause meningitis that avoid phagocytosis either by preventing deposition of complement by sialic acid on the surface or by modification of lipopolysaccharide (LPS). Haemophilus influenza expresses a mucoid polysaccharide capsule that prevents digestion by host phagocytes; a few strains resist opsonization and have become serum resistant by modification of their LPS O-antigen side chains, rendering them “invisible” to host immune defenses.

What if the nanorobot has been recognized as foreign by a white cell? As the next line of defense, medical nanorobots can directly inhibit phagocytic binding and activation. In the case of receptor-mediated binding, dansylcadaverine, amantadine, and rimantadine induce inhibition of endocytosis of complement-coated zymosan particles by human granulocytes.

These drugs block receptor-mediated endocytosis, possibly by their actions on phospholipid metabolism [23], although dansylcadaverine is not an endocytosis inhibitor in cells lacking transglutaminase activity. Cell-bound or soluble protein A produced by Staphylococcus aureus [24] attaches to the Fc region of IgG and blocks the cytophilic (cell-binding) domain of the antibody; thus the ability of IgG to act as an opsonic factor is inhibited, and opsonin-mediated ingestion of the bacteria is blocked. In the case of nonreceptor phagocytic binding, medical nanorobots could emit or expose on their surfaces chemical surfactants which would repel the lipid bilayer wall, e.g., by reducing the nanorobot’s coefficient of adhesion to very low or even negative values ([2], Section 9.2.3).

Phagocyte activation can also be directly inhibited. Several pathways of phagocytic signal transduction have been identified [25], including the activation of tyrosine kinases or serine/threonine kinase C, leading to phosphorylation of the receptors and other proteins which are recruited at the sites of phagocytosis. Monomeric GTPases of the Rho and ARF families which are engaged downstream of activated receptors, in cooperation with phosphatidylinositol 4-phosphate 5-kinase and phosphatidylinositol 3-kinase lipid modifying enzymes, can modulate locally the assembly of the submembranous actin filament system that leads to particle internalization. It may be possible for nanorobots to affirmatively influence, modulate, or even extinguish the phagocytic activation signal by physical, chemical, or other means, perhaps using GTPase or kinase inhibitors [26] such as genistein (50 microM), herbimycin (17 microM), staurosporine and trifluoperazine; in many cases there are two or more pathways that must be simultaneously inhibited, although in a few cases these pathways may share a common inhibitor. CNI-1493 is a potent and well-known macrophage deactivator or “pacifier” [27].

Acknowledgements

The author thanks C. Christopher Hook, M.D., and Stephen S. Flitman, M.D., for helpful comments on an earlier version of this paper.

Note: An article by Robert A Freitas Jr. on the potential
applications of advanced nanotechnology to dental care appeared
in the November 2000 issue of the Journal of the American Dental
Association (JADA). That article is now available from the JADA
website (http://www.ada.org/prof/pubs/jada/index.asp). To
access the article, click on the link for Archives, and choose the
options for the November 2000 issue. In the listing of the issues
contents, choose the Nanodentistry article. The article is available
as either a HTML web page or an Acrobat PDF file.

How Nanorobots Can Avoid Phagocytosis by White Cells, Part II

By Robert A. Freitas Jr.Research Scientist, Zyvex Corp.

Ingestion or phagocytosis of medical nanorobots [1] by white cells will occur in a series of well-defined steps. Normally inactive white cells are activated when they encounter a foreign particle, producing a change in metabolic activity and cell shape. During contact and recognition of the foreign particle, the phagocyte plasma membrane develops a local invagination or dimple. The particle is drawn inside and the dimple closes, often pinching off to form a small vacuole or phagosome consisting of everted cell wall membrane, trapping the particle inside the cell. The phagosome then forms a phagolysosome by merging with a lysosome, whose contents (including degradative lysozymes) are released into the smaller vacuole, attacking the enclosed foreign or denatured proteins. Afterwards the phagolysosomal vacuole may be absorbed or released to the outside at the cell’s outer surface via exocytosis, producing a large membrane flow. In cultured macrophages an amount of membrane equal to the entire surface area of the cell is replaced in ~1800 sec, and macrophages may ingest up to ~25% of their volume per hour [1].

In Part I of this paper [2], we described the initial antiphagocytic strategy for medical nanorobots which is to avoid phagocyte contact, recognition, binding and activation. In Part II, we assume that this initial strategy has either failed or has not been used, in which case the medical nanorobot has been recognized by, and become transiently bound to, a phagocyte. The best nanorobot strategies at this point are first, to inhibit phagocytic engulfment, and second, to inhibit enclosure and scission of the phagosome if engulfment has begun. Let’s look at these two approaches in more detail.

Even if a medical nanorobot has been recognized and has attached to the phagocyte outer surface (typically across a ~20 nm gap bridged by ~12 nm strands), the device can still prevent complete engulfment from taking place. Macrophages challenged with a particular type of target usually bind many more targets than they ingest [3]. Fortunately, internalization is a relatively slow process and most particles that become bound to the phagocyte surface are not ingested [3]. On rare occasions, phagocytosed particles are actually expelled.

Phagocytosis is an uptake of large particles governed by the actin-based cytoskeleton. Complement-opsonized (CO) and antibody-opsonized (AO) particles are phagocytosed differently by macrophages [4] — CO particles sink into the cell, whereas AO particles are engulfed by lamellipodia which project from the cell surface. During the ingestion of CO particles, punctate structures rich in F-actin, vinculin, alpha-actinin, paxillin, and phosphotyrosine-containing proteins are distributed over the phagosome surface [4]. These foci can be detected underneath bound CO particles within 30 seconds of cell activation, and their formation requires active protein kinase C. Complement receptor-mediated internalization requires intact microtubules and is accompanied by the accumulation of vesicles beneath the forming phagosome [4]. By contrast, during the ingestion of AO particles (Fcgamma receptor mediated phagocytosis), all proteins are uniformly distributed on or near the phagosome surface. Ingestion of AO beads is blocked by tyrosine kinase inhibitors (e.g., released from or tethered to medical nanorobots), although the phagocytosis of CO particles is not [4].

Phagocytic particle ingestion can require actin assembly and pseudopod extension, two cellular events that may coincide spatially and temporally but apparently use distinct signal transduction events or pathways [5]. Medical nanorobots that have become bound to the extracellular phagocyte surface may attempt to inhibit either or both of these signal transduction pathways.

In the first case, during actin assembly, engagement of particle-bound immunoglobulin IgG ligands by receptors for the Fc portion of IgG results in receptor aggregation and recruitment of cytosolic tyrosine kinase, especially Syk [6]. The onset of uptake is accompanied by tyrosine phosphorylation of several proteins, which persists for up to 3 minutes, is concentrated at phagocytic cups and nascent phagosomes, and is correlated with the accumulation of actin filaments. Phosphorylation of tyrosine residues occurs within immunoreceptor tyrosine activation motif (ITAM) consensus sequences found in FcgammaR subunits, which allows further recruitment and activation of Syk [6]. Syk tyrosine kinase activity is required for FcgammaR-mediated actin assembly, which is controlled by several GTPases, including Rac1 and CDC42 [6]. Rac1 and CDC42 (two Rho proteins involved in the signal transduction through the Fc receptors) are required to coordinate actin filament organization and membrane extension to form phagocytic cups, to allow particle internalization during FcR-mediated phagocytosis, and are involved in the phosphotyrosine dephosphorylation required for particle internalization [7].

Actin assembly can be inhibited by Clostridium difficile toxin B, which is a general inhibitor of Rho GTP-binding proteins [7]. Inhibition of Rac1 or CDC42 severely inhibits particle internalization but not F-actin accumulation [7]. In laboratory tests with cells, inhibition (via knockout of gene expression in a mutant line) of CDC42 function results in pedestal-like structures with foreign particles at their tips on the phagocyte surface, whereas inhibition of Rac1 results in particles bound at the surface that are enclosed within thin unfused membrane protrusions, demonstrating that Rac1 and CDC42 have distinct functions and may act cooperatively in the assembly of the phagocytic cup [7]. Phagocytic cup closure and particle internalization has also been blocked when phosphotyrosine dephosphorylation is inhibited by treatment of the phagocytic cells with phenylarsine oxide, an inhibitor of protein phosphotyrosine phosphatases [7]. Ceramide also inhibits tyrosine phosphorylation in human neutrophils [8].

In the second case, during pseudopod extension, phosphatidylinositol 3-kinase (PI3K) is recruited to the plasma membrane, triggering exocytosis from an internal membrane source, as is required for pseudopod extension [6]. (Macrophage spreading on opsonized surface is accompanied by insertion into the plasma membrane of new membrane from intracellular sources [5].) One or more isoforms of PI3K are required for maximal pseudopod extension, though not for phagocytosis per se; PI3K is required for coordinating exocytic membrane insertion and pseudopod extension [5].

Pseudopod extension may be partially inhibited using wortmannin (WM) or LY294002, which are two inhibitors of PI3K [5]. Both of these specifically inhibit phagocytosis without inhibiting Fcgamma receptor-directed actin polymerization, by preventing maximal pseudopod extension. Decreasing the size of test beads, and hence the size of pseudopod extension required for particle engulfment, de-inhibited phagocytosis (in presence of these inhibitors) by up to 80% at the very smallest submicron particle sizes. For larger (nanorobot-sized) foreign particles, phagocytosis is blocked before phagosomal closure. Both compounds also inhibit macrophage spreading on opsonized surfaces (i.e., on substrate-bound IgG) [5].

Amphiphysin II associates with early phagosomes in macrophages and participates in receptor-mediated endocytosis by recruiting the GTPase dynamin to the nascent endosome. There is a signaling cascade in which PI3K is required to recruit amphiphysin II to the phagosome, after which the amphiphysin II in turn recruits dynamin to the phagosome [9]. A modified amphiphysin II molecule with its dynamin-binding site ablated away inhibits phagocytosis at the stage of membrane extension around the bound foreign particles [9]. Both phenylbutazone and chloramphenicol also have shown an inhibitory effect on the engulfment stage of phagocytosis of yeast particles by cultured human monocytes [10].

As might be expected, bacteria already employ a wide variety of strategies to avoid engulfment when physically contacted by host phagocytes [11]. Some of these strategies could in principle be mimicked by medical nanorobots. Most commonly, many important pathogenic bacteria bear substances on their surfaces that inhibit phagocytic adsorption or engulfment. Resistance to phagocytic ingestion is usually due to an antiphagocytic component of the bacterial cell surface, such as:

Fimbriae and M Protein  fimbriae in E. coli [12], and M protein and fimbriae of Group A streptococci [12]. For example, Streptococcus pyogenes has M protein, a fibrillar surface protein whose distal end bears a negative charge that interferes with phagocytosis. Enterococci also have antiphagocytic surface proteins [13], such as M protein.

Macrophages can also bind and engulf a variety of particles in the absence of specific opsonins, a process known as nonspecific phagocytosis [14], nonopsonic phagocytosis [15], or opsonin-independent phagocytosis. Polystyrene microspheres are often used to demonstrate this. For instance, during patocytosis of hydrophobic >0.5-micron particles by phagocytes, actin-independent capping of hydrophobic polystyrene microspheres on the plasma membrane precedes actin-dependent uptake of the microspheres into the surface-connected compartments [16]. Microsphere transport from plasma membrane invaginations into spaces created by unfolding stacks of internal microvilli are inhibited by administering primaquine [16]. Studies of non-specific endocytosis and binding of liposomes by mouse peritoneal macrophages also found that particle internalization declined markedly after anchorage of the cells to polystyrene substrate [17]. Inhibitors are potentially available to medical nanorobots to halt these processes too. For example, staurosporine selectively inhibits nonspecific phagocytosis while having no effect on receptor-mediated phagocytosis [14].

What if a medical nanorobot has become partially or wholly engulfed by a phagocyte? Can the vacuole still be prevented from pinching off and separating into a free intracellular phagosome containing the nanorobot (i.e., enclosure and scission)? More research is needed, but the answer appears to be yes.

Cells normally internalize soluble ligands and small particles via endocytosis and large particles via actin-based phagocytosis. The dynamin family of GTPases mediates the membrane destabilization, constriction, fission (scission) and trafficking of endocytic vesicles from the plasma membrane, but dynamin-2 also has a role in phagocytosis by macrophages [18]. Experiments reveal that early phagosomes (vacuoles) are enriched in dynamin-2, and inactive mutant versions of this molecule, if expressed, inhibit particle internalization at the stage of membrane extension around the particle [18]. This arrest of phagocytosis resembles that seen with PI3K inhibitors, preventing the recruitment of dynamin to the site of particle binding. Dynamin is a microtubule-binding enzyme with a microtubule-activated GTPase activity; phosphorylation engages its activity. Dynamin can interact with the actin cytoskeleton to regulate actin reorganization and subsequently cell shape [19].

Observations suggest that dynamin mediates scission from the plasma membrane of both clathrin-coated pits and caveolae during distinct endocytic processes [20]. For example, dynamin-1 is a 100 kD GTPase involved in scission of endocytic vesicles from the plasma membrane. It is present in solution as tetramer, undergoes self-assembly (following its recruitment to coated pits) to form higher-order oligomers that resemble collars around the necks of nascent coated buds. GTPase hydrolysis by dynamin in these collars is thought to accompany the pinching off of endocytic vesicles — dynamin may use GTPase hydrolysis physically to drive vesiculation, or may act as a classical G protein switch, or both [21]. (Purified dynamin readily self-assembles into rings or spirals, suggesting that it probably wraps around the necks of budding vesicles and squeezes, pinching them off [22]; the large GTPase dynamin is a mechanoenzyme [23].) Different dynamin isoforms may be localized to distinct cellular compartments but provide similar scission functions during the biogenesis of nascent cytoplasmic vesicles [20]. Once again, inhibitory tools that might be employed by medical nanorobots are potentially available. For example, anti-dynamin antibodies have been used to specifically inhibit dynamin function in cultured mammalian epithelial cells, inhibiting cellular uptake of external particles in these cells [20], and to inhibit clathrin-mediated endocytosis in hepatocytes [24]. Ca++ inhibits both dynamin I GTPase [25] and dynamin II GTPase [26] and may also serve as vesiculation inhibitors for fully engulfed medical nanorobots. Alternatively, butanedione monoxime, a class II myosin inhibitor, has been shown to prevent the purse-string-like contraction that closes phagosomes while not inhibiting the initial pseudopod extension [27].

One important remaining practical question is how the various inhibitory methods that we have identified will be specifically expressed in each class of medical nanorobots, but this is an issue for another time.

Acknowledgements

The author thanks Stephen S. Flitman, M.D., and C. Christopher Hook, M.D., for helpful comments on an earlier version of this paper.

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